U.S. patent number 6,926,670 [Application Number 10/054,330] was granted by the patent office on 2005-08-09 for wireless mems capacitive sensor for physiologic parameter measurement.
This patent grant is currently assigned to Integrated Sensing Systems, Inc.. Invention is credited to Sonbol Massoud-Ansari, Nader Najafi, Collin A. Rich, Matthew Z. Straayer, Yafan Zhang.
United States Patent |
6,926,670 |
Rich , et al. |
August 9, 2005 |
**Please see images for:
( PTAB Trial Certificate ) ** |
Wireless MEMS capacitive sensor for physiologic parameter
measurement
Abstract
The present invention relates to an implantable microfabricated
sensor device and system for measuring a physiologic parameter of
interest within a patient. The implantable device is micro
electromechanical system (MEMS) device and includes a substrate
having an integrated inductor and at least one sensor formed
thereon. A plurality of conductive paths electrically connect the
integrated inductor with the sensor. Cooperatively, the integrated
inductor, sensor and conductive paths defining an LC tank
resonator.
Inventors: |
Rich; Collin A. (Ypsilanti,
MI), Zhang; Yafan (Plymouth, MI), Najafi; Nader (Ann
Arbor, MI), Straayer; Matthew Z. (Ann Arbor, MI),
Massoud-Ansari; Sonbol (Ann Arbor, MI) |
Assignee: |
Integrated Sensing Systems,
Inc. (Ypsilanti, MI)
|
Family
ID: |
26949776 |
Appl.
No.: |
10/054,330 |
Filed: |
January 22, 2002 |
Current U.S.
Class: |
600/459 |
Current CPC
Class: |
A61B
5/6876 (20130101); A61B 5/6852 (20130101); A61B
5/6862 (20130101); A61B 5/0215 (20130101); A61B
5/02028 (20130101); A61B 5/6882 (20130101); A61B
5/4839 (20130101); A61B 5/14539 (20130101); A61B
5/0295 (20130101); A61B 5/14546 (20130101); A61B
5/0031 (20130101); A61B 5/027 (20130101); A61B
5/14532 (20130101); A61B 2562/028 (20130101); A61B
5/1473 (20130101) |
Current International
Class: |
A61B
5/0215 (20060101); A61B 5/00 (20060101); A61B
5/027 (20060101); A61B 5/026 (20060101); A61B
008/14 () |
Field of
Search: |
;600/437-472 ;73/625-633
;367/7,11,130,138 ;607/36-38,1,2,60 ;128/916 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
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0812016 |
|
Dec 1997 |
|
EP |
|
WO00/19888 |
|
Apr 2000 |
|
WO |
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WO00/30534 |
|
Jun 2000 |
|
WO |
|
Other References
A Passive Wireless Integrated Humidity Sensor, Timothy Harpster et
al. 2001, pp. 553-557. .
Electrodeposited Copper Inductors for Intraocular pressure
Telemetry; R. Puers et al.; 2001 pp. 124-129. .
Hermetically Sealed Inductor-Capacitor (LC) Resonator for Remote
Pressure Monitoring; Eun-Chul Park et al.; Sep. 8, 1998; pp.
7124-7128. .
Micromachined Planar Inductors on Silicon Wafers for MEMS
Applications; Chong H. Ahn et al..
|
Primary Examiner: Imam; Ali
Attorney, Agent or Firm: Brinks Hofer Gilson & Lione
Parent Case Text
CROSS REFERENCE TO RELATED-APPLICATION
This application claims priority to prior U.S. provisional
application No. 60/263,327 (filed Jan. 22, 2001) and U.S.
provisional application No. 60/278,634 (filed Mar. 26, 2001).
Claims
What is claimed is:
1. An implantable microfabricated sensor device for measuring a
physiologic parameter of interest within a patient, said sensor
comprising: an implantable sensing device, said sensing device
being a micro electromechanical system (MEMS) comprising a
substrate, an integrated inductor formed on the substrate, at least
one sensor responsive to the physiologic parameters and being
formed at least in part on the substrate, a plurality of conductive
paths electrically connecting said integrated inductor with said
sensor, said integrated inductor, said sensor and said conductive
paths cooperatively defining an LC tank resonator.
2. The sensor device of claim 1 wherein said sensor is a capacitive
sensor having a fixed electrode and a moveable electrode.
3. The sensor device of claim 2 wherein said fixed electrode is
formed on said substrate.
4. The sensor device of claim 2 wherein said sensor is a pressure
sensor.
5. The sensor device of claim 2 wherein said sensor is a
temperature sensor.
6. The sensor device of claim 2 wherein said sensor is a chemical
sensor.
7. The sensor device of claim 1 wherein said integrated inductor
includes a magnetic core and a winding comprised of a conductive
material about said magnetic core.
8. The sensor device of claim 7 wherein said magnetic core includes
a plate member formed on a first face of said substrate.
9. The sensor device of claim 8 wherein said magnetic core further
includes a second plate member, said second plate member being
formed on a second face of said substrate and located generally
opposite of said first plate member.
10. The sensor device of claim 9 further comprising a post
extending through said substrate and connecting said first plate to
said second plate.
11. The sensor device of claim 8 wherein said winding is formed
within said first plate.
12. The sensor device of claim 8 further comprising a cap layer
formed over said plate member.
13. The sensor device of claim 12 wherein said cap layer includes a
portion defining a moveable electrode of said sensor.
14. The sensor device of claim 12 wherein said cap layer is
conductive.
15. The sensor device of claim 12 wherein said cap layer is doped
silicon.
16. The sensor device of claim 7 wherein said magnetic core
includes first and second plate members connected to one another by
post.
17. The sensor device of claim 16 wherein said windings are about
said post.
18. The sensor device of claim 17 wherein said windings are about
said post and adjacent to said first plate.
19. The sensor device of claim 1 further comprising active
circuitry being formed in said sensing device.
20. The sensor device of claim 19 wherein said active circuitry is
formed within a cap layer formed over said integrated inductor.
21. The sensor device of claim 1 wherein said sensor device is
wireless.
22. The sensor device of claim 1 wherein said sensing device is
monolithic.
23. The sensor device of claim 1 further comprising at least two
sensors.
24. The sensor device of claim 23 wherein said two sensors sense
the same physiologic parameter.
25. The sensor device of claim 23 wherein said two sensors sense
different physiologic parameters.
26. The sensor device of claim 1 wherein said sensor is a
capacitive sensor including a fixed electrode and a moveable
electrode, said fixed and moveable electrodes defining a chamber
therebetween, said chamber being in fluid communication with a
displacement cavity.
27. The sensor device of claim 26 wherein said displacement cavity
is defined within said substrate.
28. The sensor device of claim 1 wherein said sensor is a
capacitive sensor having a fixed electrode and a moveable
electrode, said fixed and moveable electrodes being electrically
coupled by first and second traces to said integrated inductor,
said first and second traces being electrically isolated from one
another.
29. The sensor device of claim 28 wherein said traces are isolated
by a dielectric layer therebetween.
30. The sensor device of claim 28 wherein said traces are isolated
by a p-n junction structure.
31. The sensor device of claim 1 as part of a sensing system
further comprising a non-implantable readout device, said readout
device including a second inductor adapted to magnetically couple
with said integrated inductor to read changes in said LC tank
resonator as a result of said sensor sensing the physiologic
parameter of interest.
Description
BACKGROUND OF THE INVENTION
Field of the Invention
The present invention generally relates to the field of MEMS
(micro-electromechanical systems) sensors and more specifically to
a wireless MEMS capacitive sensor for implantation into the body of
a patient to measure one or more physiologic parameters.
A number of different biologic parameters are strong candidates for
continuous monitoring. These parameters include, but are not
limited to blood pressure, blood flow, intracranial pressure,
intraocular pressure, glucose levels, etc. Wired sensors, if used
have certain inherent limitations because of the passage of wires
(or other communication "tethers") through the cutaneous layer.
Some limitations include the risks of physical injury and infection
to the patient. Another risk is damage to the device if the wires
(the communication link) experience excessive pulling forces and
separate from the device itself. Wireless sensors are therefore
highly desirable for biologic applications.
A number of proposed schemes for wireless communication rely on
magnetic coupling between an inductor coil associated with the
implanted device and a separate, external "readout" coil. For
example, one method of wireless communication (well-known to those
knowledgeable in the art) is that of the LC (inductor-capacitor)
tank resonator. In such a device, a series-parallel connection of a
capacitor and inductor has a specific resonant frequency, expressed
as 1/√LC, which can be detected from the impedance of the circuit.
If one element of the inductor-capacitor pair varies with some
physical parameter (e.g. pressure), while the other element remains
at a known value, the physical parameter may be determined from the
resonant frequency. For example, if the capacitance corresponds to
a capacitive pressure sensor, the capacitance may be
back-calculated from the resonant frequency and the sensed pressure
may then be deduced from the capacitance by means of a calibrated
pressure-capacitance transfer function.
The impedance of an LC tank resonator may be measured directly or
it may also be determined indirectly from the impedance of a
separate readout coil that is magnetically coupled to the internal
coil. The latter case is most useful for biologic applications
since the sensing device may be subcutaneously implanted, while the
readout coil may be located external to the patient, but in a
location that allows magnetic coupling between the implanted
sensing device and readout coil. It is possible for the readout
coil (or coils) to simultaneously excite the resonator of the
implanted device and sense the reflected back impedance.
Consequently, this architecture has the substantial advantage of
requiring no internal power source, which greatly improves its
prospects for long-term implantation (e.g. decades to a human
lifetime).
Such devices have been proposed in various forms for many
applications. Chubbuck (U.S. Pat. No. 4,026,276), Bullara (U.S.
Pat. No. 4,127,110), and Dunphy (U.S. Pat. No. 3,958,558) disclose
various devices initially intended for hydrocephalus applications
(but also amenable to others) that use LC resonant circuits. The
'276, '110, and '558 patents, although feasible, do not take
advantage of recent advances in silicon (or similar)
microfabrication technologies. Kensey (U.S. Pat. No. 6,015,386)
discloses an implantable device for measuring blood pressure in a
vessel of the wrist. This device must be "assembled" around the
vessel being monitored such that it fully encompasses the vessel,
which may not be feasible in many cases. In another application,
Frenkel (U.S. Pat. No. 5,005,577) describes an implantable lens for
monitoring intraocular pressure. Such a device would be
advantageous for monitoring elevated eye pressures (as is usually
the case for glaucoma patients); however, the requirement that the
eye's crystalline lens be replaced will likely limit the general
acceptance of this device.
In addition to the aforementioned applications that specify LC
resonant circuits, other applications would also benefit greatly
from such wireless sensing. Han, et al. (U.S. Pat. No. 6,268,161)
describe a wireless implantable glucose (or other chemical) sensor
that employs a pressure sensor as an intermediate transducer (in
conjunction with a hydrogel) from the chemical into the electrical
domain.
The treatment of cardiovascular diseases such as Chronic Heart
Failure (CHF) can be greatly improved through continuous and/or
intermittent monitoring of various pressures and/or flows in the
heart and associated vasculature. Porat (U.S. Pat. No. 6,277,078),
Eigler (U.S. Pat. No. 6,328,699), and Carney (U.S. Pat. No.
5,368,040) each teach different modes of monitoring heart
performance using wireless implantable sensors. In every case,
however, what is described is a general scheme of monitoring the
heart. The existence of a method to construct a sensor with
sufficient size, long-term fidelity, stability, telemetry range,
and biocompatibility is noticeably absent in each case, being
instead simply assumed. Eigler, et al., come closest to describing
a specific device structure although they disregard the baseline
and sensitivity drift issues that must be addressed in a long-term
implant. Applications for wireless sensors located in a stent
(e.g., U.S. Pat. No. 6,053,873 by Govari) have also been taught,
although little acknowledgement is made of the difficulty in
fabricating a pressure sensor with telemetry means sufficiently
small to incorporate into a stent.
Closed-loop drug delivery systems, such as that of Feingold (U.S.
Pat. No. 4,871,351) have likewise been taught. As with others,
Feingold overlooks the difficulty in fabricating sensors that meet
the performance requirements needed for long-term implantation.
In nearly all of the aforementioned cases, the disclosed devices
require a complex electromechanical assembly with many dissimilar
materials, which will result in significant temperature- and
aging-induced drift over time. Such assemblies may also be too
large for many desirable applications, including intraocular
pressure monitoring and/or pediatric applications. Finally, complex
assembly processes will make such devices prohibitively expensive
to manufacture for widespread use.
As an alternative to conventionally fabricated devices,
microfabricated sensors have also been proposed. One such device is
taught by Darrow (U.S. Pat. No. 6,201,980). Others are reported in
the literature (see, e.g. Park, et al., Jpn. J. Appl. Phys., 37
(1998), pp. 7124-7128; Puers, et al., J. Micromech. Microeng. 10
(2000), pp. 124-129; Harpster et al., Proc. 14.sup.th IEEE Int'l.
Conf. Microelectromech. Sys. (2001), pp. 553-557).
Past efforts to develop wireless sensors have separately located
the sensor and inductor and have been limited to implant-readout
separation distances of 1-2 cm at most, rendering them impractical
for implantation much deeper than immediately below the cutaneous
layer. This eliminates from consideration wireless sensing
applications, such as heart ventricle pressure monitoring or
intracranial pressure monitoring, that inherently require
separation distances in the range of 5-10 cm. In the present
state-of-the-art, several factors have contributed to this
limitation on the separation distance including 1) signal
attenuation due to intervening tissue, 2) suboptimal design for
magnetic coupling efficiency; and 3) high internal energy losses in
the implanted device.
In view of the above and other limitations on the prior art, it is
apparent that there exists a need for an improved wireless MEMS
sensor system capable of overcoming the limitations of the prior
art and optimized for signal fidelity, transmission distance and
manufacturability. It is therefore an object of the present
invention is to provide a wireless MEMS sensor system in which the
sensing device is adapted for implantation within the body of
patient.
A further object of this invention is to provide a wireless MEMS
sensor system in which the separation distance between the sensing
device and the readout device is greater than 2 cm, thereby
allowing for deeper implantation of the sensing device within the
body of a patient.
Still another object of the present invention is to provide a
wireless MEMS sensor system in which the sensing device utilizes an
integrated inductor, an inductor microfabricated with the sensor
itself.
It is also an object of this invention to provide a wireless MEMS
sensor system in which the sensing device is batteryless.
A further object of the present invention is to provide a wireless
MEMS sensor system.
BRIEF SUMMARY OF THE INVENTION
In overcoming the limitations of the prior art and achieving the
above objects, the present invention provides for a wireless MEMS
sensor for implantation into the body of a patient and which
permits implantation at depths greater than 2 cm while still
readily allowing for reading of the signals from the implanted
portion by an external readout device.
In achieving the above, the present invention provides a MEMS
sensor system having an implantable unit and a non-implantable
unit. The implantable unit is microfabricated utilizing common
microfabricating techniques to provide a monolithic device, a
device where all components are located on the same chip. The
implanted device includes a substrate on which is formed a
capacitive sensor. The fixed electrode of the capacitive sensor may
formed on the substrate itself, while the moveable electrode of the
capacitive sensor is formed as part of a highly doped silicon layer
on top of the substrate. Being highly doped, the silicon layer
itself operates as the conductive path for the moveable electrode.
A separate conductive path is provided on the substrate for the
fixed electrode.
In addition to the capacitive sensor, the implanted sensing device
includes an integrally formed inductor. The integral inductor
includes a magnetic core having at least one plate and a coil
defining a plurality of turns about the core. One end of the coil
is coupled to the conductive lead connected with the fixed
electrode while the other end of the coil is electrically coupled
to the highly doped silicon layer, thereby utilizing the silicon
layer as the conductive path to the moveable electrode.
In order to optimize the operation of the inductor and to permit
greater implantation depths, a novel construction is additionally
provided for the magnetic core. In general, the optimized magnetic
core utilizes a pair of plates formed on opposing sides of the
substrate and interconnected by a post extending through the
substrate. The windings of the coil, in this instance, are provided
about the post.
The external readout device of the present system also includes a
coil and various suitable associated components, as well known in
the field, to enable a determination of the pressure or other
physiologic parameter being sensed by the implanted sensing device.
The external readout device may similarly be utilized to power the
implanted sensing device and as such the implanted sensing device
is wireless.
Integrally formed on the implanted device and microfabricated
therewith, may be additionally be active circuitry for use in
conjunction with capacitive sensor. Locating this circuitry as near
as possible to the capacitive sensor minimizes noise and other
factors which could lead to a degradation in the received signal
and the sensed measured physiologic parameter. As such, the active
circuitry may be integrally microfabricated in the highly doped
silicon layer mentioned above.
Further object and advantages of the present invention will become
apparent to those skilled in the art from a review of the drawings
in connection with the following description and dependent
claims.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a schematic illustration of a wireless MEMS sensor system
according the principles of the present invention;
FIG. 2 is a graphical illustration of impedance magnitude and phase
angle near resonance, as sensed through a readout coil;
FIG. 3 is a cross-sectional representation of a sensing device
embodying the principles of the present invention.
FIGS. 4A and 4B are schematic illustrations of the magnetic field
distribution with FIG. 4A illustrating the magnetic field
distribution of prior art devices and with FIG. 4B illustrating the
magnetic field distribution for a sensing device having a magnetic
core embodying the principles of the present invention;
FIG. 5 is an enlarged cross-sectional view of the diaphragm portion
of FIG. 3 operating in what is herein referred to as a "proximity"
mode;
FIG. 6 is a cross-sectional view similar to that seen in FIG. 5
illustrating, however, the diaphragm operating in what is herein
referred to as a "touch" mode;
FIG. 7 is a capacitance versus pressure curve in the proximity and
touch modes of operation;
FIG. 8 is a top plane view of a second embodiment of the main
electrode in the capacitive sensor portion of the implanted sensing
device according to the principles of the present invention;
FIG. 9 is a diagrammatic illustration of one scheme for providing
electrically isolated paths for the connections and electrodes of
the capacitive sensor portion;
FIG. 10 is a diagrammatic illustration of another scheme for
electrically isolating the conductive paths for the connections and
contacts of the capacitive sensor portion;
FIG. 11 is a cross-sectional view, generally similar to that seen
in FIG. 3, further incorporating active circuitry into the sensing
device;
FIG. 12 is a block diagram illustrating one possible circuit
implementation of the active circuitry when incorporated into the
sensing device of the present wireless MEMS sensing system;
FIG. 13 illustrates one method of mounting, within the body of a
patient, a sensing device embodying the principles of the presents
invention;
FIG. 14 illustrates a second embodiment by which a sensing device
embodying the principles of the present invention may be secured to
tissues within the body of a patient
FIGS. 15 and 16 are diagrammatic illustrations of different
embodiments for locating a sensing device according to the
principles of the present invention, within a vessel in the body of
a patient;
FIG. 17 illustrates a sensing device, according to the principles
of the present invention, encapsulated in a material yielding a
pellet-like profile for implantation into the tissues in the body
of a patient;
FIG. 18 illustrates a sensing device according to the principles of
the present invention being located within the electrode tip of an
implantable stimulation lead, such as that used for cardiac
pacing;
FIG. 19 illustrates a plurality of sensing devices according to the
present invention located within a catheter and utilized to
calculate various physiologic parameters within a vessel within the
body of a patient;
FIG. 20 is a schematic illustration of multiple sensors being used
to measure performance of a component in the body or a device
mounted within the body of a patient;
FIG. 21 illustrates a sensing device according to the principles of
the present invention being utilized to measure pressure externally
through a vessel wall;
FIG. 22 illustrates a portion of a further embodiment of the
present invention in which the pressure sensing features of the
sensing device have been augmented over or replaced with a
structure allowing a parameter other than pressure to be
sensed;
FIG. 23 is schematic perspective view, with portions enlarged,
illustrating an alternative embodiment for sensing according to the
principles of the present invention; and
FIG. 24 is an embodiment generally similar to that seen in FIG. 23
for sensing according to the principles of the present
invention.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
In order to provide for battery-less, wireless physiologic
parameter sensing over significant distances greater than 2 cm
(e.g. 10 cm), the present invention provides a wireless MEMS
sensing system, generally designated at 10 and seen schematically
in FIG. 1. The system 10 includes a microfabricated implantable
sensing device 12, optimized for coupling with an external readout
device 14. The sensing device 12 is provided with an integrated
inductor 16 that is conductive to the integration of transducers
and/or other components necessary to construct the wireless sensing
system 10. As an example, the preferred embodiment integrates a
capacitive pressure sensor 18 into a common substrate 20 with the
integrated inductor 16. A second inductor 24, in the readout device
14, couples magnetically 26 with the integrated inductor 16 of the
sensing device 12.
The readout device 14 is constructed according to techniques well
known in the industry and in the sensing field in general. As such,
the readout device 14 is not illustrated or described in great
detail. It is noted, however, that the readout device 14 may be
included, in addition to its inductor 24, signal conditioning,
control and analysis circuitry and software, display and other
hardware and may be a stand alone unit or may be connected to a
personal computer (PC) or other computer controlled device.
The magnetic coupling 26 seen in FIG. 1 allows the impedance of the
LC tank circuit 22 to be sensed by the readout device 14. The
typical impedance magnitude 28 and phase angle 30 near resonance
32, as sensed through the readout coil 14, is seen in FIG. 2.
Real-time measurement and analysis of this impedance and changes
therein allows the sensed pressure to be determined as previously
mentioned.
Referring now to FIG. 3, a cross section of a preferred embodiment
of the sensing device 12 is illustrated therein. The sensing device
12 includes a main substrate 34 (preferably 7740 Pyrex glass)
formed and located within recessed regions of the substrate 34 are
those structures forming the integrated inductor 16. The integrated
inductor 16 is seen to include a magnetic core 33 defined by a top
plate 36, a bottom plate 38 and a post 40 connecting the top plate
36 to the bottom plate 38 and being continuous through the
substrate 34. The plates 36 and 38 and the post 40 are preferably
constructed of the same material, a ferromagnetic material and are
monolithic. The integrated conductor 16 additionally includes a
coil 42, preferably composed of copper or other high-conductivity
material, successive turns of which surround the post 40 of the
magnetic core 33.
In FIG. 3, the coil 42 is seen as being recessed into the top plate
36. The coil 42 may additionally be planar or layered and
preferably wraps as tightly as possible about the post 40. If the
material of the coil 42 has a high electrical resistance relative
to the material of the core 33, (as in a copper coil and NiZn
ferrite core system) the core 33, and specifically the top plate 36
may be directly deposited on top of the coil 42 without need for a
intermediate insulating layer. If the electrical resistance of the
coil material relative to the coil material is not high, an
intermediate insulating layer must be included between the
successive turns of the coil 42 and the core 33.
Top and bottom cap layers 44 and 46 respectively, are provided over
upper and lower faces 48 and 50 of the substrate 20 and over the
top and bottom plates 36 and 38 of the magnetic core 33. To
accommodate any portions of the magnetic core 33 that extend
significantly above or below the upper and lower faces 48 and 50 of
the substrate 20, the cap layers 44 and 46 may be provided with
recesses 52 and 54, respectively. Preferably, the cap layers 44 and
46 are of monocrystalline silicon. Other preferred materials
include polycrystalline silicon, epitaxially deposited silicon,
ceramics, glass, plastics, or other materials that can be bonded to
lower substrate and/or are suitable for fabrication of the sensor
diaphragm. In lieu of a monolithic cap layer, several sub-pieces
may be fabricated at separate process steps, together forming a
complete cap layer after processing is finished.
The coupling effectiveness of the integrated inductor 16 is a
function of the magnetic flux enclosed by the windings of the coil
42; therefore the coupling is greatest if the structure of the
integrated inductor 16 maximizes the flux encompassed by all of the
winding loops. FIG. 4A shows schematically the magnetic field
distribution 56 in a known inductor structure having a single core
layer 58 and associated windings 60. Schematically shown in FIG. 4b
is the magnetic field distribution 62 for an inductor structure 16'
having upper and lower plates 36' and 38', connected by a post 40'
about which windings of a coil 42' are located, as generally seen
in the present invention. The design of the present invention
optimizes the inductor geometry for maximum field coupling. Placing
the plates 36 and 38 on opposite sides of the substrate 20, as in
FIG. 3, increases the plate-to-plate spacing. The increased plate
spacing creates a localized path of least resistance for the
free-space magnetic field of an external readout coil, causing the
magnetic field to preferentially pass through the post 40 of the
integrated inductor's magnetic core 33. This increases device
effectiveness since the coupling efficiency between the sensor and
a readout unit increases with the total magnetic flux encompassed
by the windings of the inductor. A greater coupling efficiency
increases the maximum separation distance between the sensor and a
readout unit.
The materials used to form the integrated inductor 16 should be
chosen and/or processed to maximize the above mentioned effect and
to minimize drift in the inductance value across time, temperature,
package stress, and other potentially uncontrolled parameters. A
high-permeability material such as NiZn ferrite is used to maximize
this effect on the magnetic field and to minimize drift. Other
preferred materials include nickel, ferrite, permalloy, or similar
ferrite composites.
To the right of the integrated inductor 16 seen in FIG. 3 is the
capacitive pressure sensor 18. The capacitive pressure sensor 18
may be constructed in many forms commonly know to those familiar
with the art. In the illustrated embodiment, the upper cap layer 44
is formed to define a diaphragm 64. The diaphragm 64 constitutes
and may also be referred to as the moveable electrode of the
pressure sensor 18. The fixed electrode 66 of the pressure sensor
18 is defined by a conductive layer formed on the upper face 48 of
the substrate 20, in a position immediately below the moveable
electrode or diaphragm 64. If desired, a conductive layer may
additionally be located on the underside of the moveable electrode
64. To prevent shorting between the upper electrode 64 (as defined
by either the diaphragm itself or the diaphragm and the conductive
layer 68) and the lower electrode 66, one or both of the electrodes
64 and 66 may be provided with a thin dielectric layer (preferably
less than 1000 .ANG.) deposited thereon.
To improve performance of the capacitive pressure sensor 18, as
seen in FIG. 8, one or more secondary electrodes designated at 70
may be located about the fixed electrode 66 near the projected edge
of the diaphragm 64 where pressure induced deflection of the
diaphragm 64 is minimal. The secondary electrodes 70 experience all
of the capacitance-effecting phenomena seen by the main electrode
66, with the exception of any pressure-induced phenomena. The
secondary electrodes 70, as such, operate as reference electrodes
and by subtracting the secondary electrodes' capacitive measurement
from the capacitive measurement of the main electrode 66, most or
all non-pressure-induced capacitance changes (signal drift) may be
filtered out. Examples as sources of signal drift, that may be
filtered out by this method, include thermally induced physical
changes and parasitics resulting from an environment with changing
dielectric constant, such as insertion of the sensor into tissue.
In a preferred embodiment, the secondary (or reference) electrodes
70 would require an additional coil, similar to construction of the
previously mentioned coil 42 to form a separate LC tank circuit. It
is noted, that both coils may, however, share the same core post
40.
Under normal operation, pressure applied to the exterior or top
surface of the capacitive pressure sensor 18 causes the diaphragm
64 (or at least the center portions thereof) to deflect downward
toward the fixed electrode 66. Because of the change in distance
between the fixed electrode 66 and the moveable electrode 64, a
corresponding change will occur in the capacitance between the two
electrodes. The applied pressure is therefore translated into a
capacitance. With this in mind, it is seen that the capacitance
pressure sensor 18 may be operated in either of two modes.
A first mode, hereinafter referred to as the "proximity" mode, is
generally seen in FIG. 5. In this mode of operation, the starting
gap between the fixed electrode 66 and the moveable electrode 64,
as well as the material and physical parameters for the diaphragm
64 itself, are chosen such that the fixed electrode 66 and the
moveable electrode 64will be spaced apart from one another over the
entire operating pressure range of the sensor 18. For the standard
equation of parallel plate capacitance, C=.di-elect cons.A/d, the
plate separation d will vary with the applied pressure, while the
plate area A and the permittivity .di-elect cons. remain
constant.
In the touch mode of operation, generally seen in FIG. 6, the
geometry (e.g., initial gap spacing between the fixed electrode 66
and the moveable electrode 64) as well as the material and physical
parameters of the diaphragm 64 itself, are chosen such that the
fixed electrode 66 and the moveable electrode 64 will progressively
touch each other over the operating pressure range of the sensor
18. Accordingly, the area 72 of the fixed electrode 66 and the
moveable electrode 64 in contact with each other will vary with the
applied pressure. In the touch mode of operation, the dominant
capacitance is the capacitance of the regions of the fixed
electrode 66 and the moveable electrode 64 in contact with one
another (if the dielectric coating 74 is thin compared to the total
gap thickness, thereby yielding a relatively small effective plate
separation distance d). In the capacitance equation mentioned
above, plate separation d and permittivity .di-elect cons. will
remain constant (at approximately that of the dielectric thickness)
while the plate contact area A varies with the applied
pressure.
In the graph of FIG. 7, capacitance-pressure relationship in the
proximity and touch modes, respectively designated at 76 and 78,
are seen. From a practical standpoint, the operational mode may be
chosen based upon sensitivity, linearity, and dynamic range
requirements. The touch mode typically yields higher sensitivity
with a more linear output, but involves mechanical contact between
surfaces and therefore requires a careful choice of the materials
to avoid wear induced changes in performance of the pressure sensor
18.
To permit the innermost turn of the coil 42 to be electrically
connected to the moveable electrode 66, a post 80 (formed integral
with the substrate 20) extends upward through the top plate 36 and
a conductive trace 82 runs up the side of the post 80. The trace 82
begins at the innermost turn of the coil 42 and proceeds to a point
where the trace 82 makes electrical contact with the upper cap
layer 44. Preferably of monocrystalline silicon and highly doped to
be conductive, the upper cap layer 44 serves as the electrical
connection between the trace 82 and moveable electrode 64. If the
upper cap layer 44 is not conductive, an additional conductive
trace along the upper cap layer 44 to the moveable electrode 64
will be utilized. The outermost turn of the coil 42 is connected by
an electrical trace 84. Where the upper cap layer 44 is conductive,
a dielectric layer 86 insulates the trace 84 from the upper cap
layer 44. Alternatively, a p-n junction structure (as further
described below) could be used.
It is noted that the inner and outer turns of the coil 42 may be
alternatively connected respectively to the fixed electrode 66 and
the moveable electrode 64, thereby reversing the polarity of the LC
tank circuit 22 if desired. Additionally, the particular paths
between the coil 42 and the electrodes 66 and 64 may also be varied
(e.g., such that both are included on the substrate 20) as best
suited by the fabrication process. In all cases, the resistance of
the electrical path through the traces 82, 84 and the upper cap
layer 44 (if used) should be minimized.
The upper and lower cap layers 44 and 46 are bonded to the
substrate 20 preferably via a hermetic sealing process.
Alternatively, a post-bond coating of the entire sensing device 12
may be used to establish hermeticity. In either situation, steps
are taken to minimize the residual gas pressure within the sensing
device 12 after a hermetic seal is established. Once the initial
hermetic seal is achieved, gas may be trapped in the interior of
the sensing device 12 due to continued outgassing of the interior
surfaces and/or the bonded regions. Gas pressure of the residual
gas will increase within the interior chamber 90 of the pressure
sensor 18 as the diaphragm 64 deflects during normal operation.
This residual gas may effect the overall sensitivity of the
pressure sensor 18 by effectively increasing the spring constant of
the diaphragm 64. Additionally, the residual gas will expand and/or
contract with changes in the temperature of the sensing device 12
itself, causing signal drift.
To compensate for the various negative effects of any residual gas,
the pressure sensor 18 is provided with a displacement cavity 88.
This displacement cavity 88 is generally seen in FIG. 3 and is in
communication either directly or through a small connecting channel
with the interior chamber 90 of the pressure sensor 18, defined
between the diaphragm 64 and the fixed electrode 66. The
displacement cavity 88 is sized such that the total internal sensor
volume, the combined volume of the displacement cavity 88 and the
interior chamber 90, varies minimally with deflection of the
diaphragm 64 over its operational range of displacement. By
minimizing the overall change in volume with deflection of the
diaphragm 64, the effect of the residual gasses are minimized and
substantially eliminated. In the preferred embodiment, the volume
of the displacement cavity 88 is approximately ten times greater
than the volume of the chamber 90. To further reduce temperature
induced drift and to increase the sensitivity of the device 12,
lower pressures within the internal volume 90 should be used.
In addition to the preferred embodiment, other configurations for
the sensing device 12 are possible. Depending on the relative sizes
of the diaphragm 64 and coil 42, the diaphragm 64 may be located
within, above, or below the turns of the coil 42, as well as off to
one end or side of the device 12 as seen in FIG. 3. The post 40
and/or one of the plates 36 or 38 of the magnetic core 33, may be
omitted to simplify fabricating. However, this would be to the
detriment of performance. Alternate lead transfer schemes may be
used instead of the disclosed traces 82 and 84 that connect the
coil 42 to the sensor 18. More or fewer wafer layers may be used to
adapt manufacturing processing to available technologies. For
example, the entire magnetic core 33 could be formed on the top
side of the substrate 20, thereby eliminating the need for lower
cap layer 46. Multiple coil layers could also be implemented to
increase the coil turn count. Finally, the overall shape of the
device 10 may be square, round, oval, or another shape.
To isolate the internal volume of the pressure sensor 18 from the
internal volume of the integrated inductor 16, a hermetic lead
transfer can be provided as a substitute for the dielectric layer
86. A hermetic lead transfer would eliminate outgassing from the
inductor coil 42 and magnetic core 33 as a source of drift for the
pressure sensor 18, thereby improving long-term stability. The
hermetic lead transfer may be accomplished by any of several means
that provide a sealed and electrically isolated conductive path.
One example, of a mechanism for achieving a sealed and electrically
isolated conductive path is through the use of a p-n junction
structure 92 in the sensor 18'. This is illustrated in FIG. 9. The
p-n junction structure 92 (with p-material forming the diaphragm)
forms an electrically isolated path in a silicon layer and provides
for electrical contact between a fixed electrode 66' and a lead
trace 94 but not from the fixed electrode 66' to the diaphragm
66'.
In another alternative construction, a separate polysilicon layer
96 forms a conductive path to a fixed electrode 66". The conductive
layer 96 is insulated, by a separate insulating layer 98, from the
doped silicon rim 100 of the sensor 18".
An alternative embodiment of the present sensing device, designated
as 12", includes active circuitry for immediate processing of the
data including logging, error correction, encoding, analysis,
multiplexing of multiple sensor inputs, etc. Since the sensing
device 12" of this embodiment, seen in FIG. 11, includes numerous
structures which are the same or identical to the structures seen
in the embodiment illustrated in FIG. 3, like structures are
accordingly provided with like designations and are not
repetitively discussed. Reference should therefore be accordingly
made to the preceding sections of this description where those
structures are discussed in connection with FIG. 3.
The block diagram of FIG. 12 illustrates one possible circuit
implementation for the active circuitry 102 seen in FIG. 11. In the
illustrated configuration, the integrated inductor 16 serves as an
antenna for RF telemetry with the external readout device 14. Using
RF modulation schemes well know to those skilled in the art, the RF
magnetic field 26 transmitted from the device 14 provides both data
communication and necessary power to the circuitry 102. The
received energy across inductor 16 is rectified and stored
temporarily in an onboard capacitor or power supply designated at
block 104. The input decoder 103 may receive digital data
pertaining to short or long term memory or real time clock signals,
and may transfer this information to the control logic 107. The
front end conditioning circuitry 109 converts an analog sensor
signal into a form that is encoded and amplified by the output
driver 105. The integrated inductor 16 then serves to transmit the
RF signal back to the external readout device 14, where the
information can be processed, stored, or displayed. The many
variations for circuit implementations of the rectifier of 104,
modulation and coding schemes encompassing blocks 103 and 105,
analog circuitry 109 and needed control logic 103 will be
appreciated.
A key issue for sensing physiologic parameters in medical
applications is that the sensor must be biocompatible.
Biocompatibility involves two issues: the effect of the sensor on
the body (toxicity), and the effect of the body on the sensor
(corrosion rate). While the fabrication of the substrate 20 of
Pyrex glass, as described in connection with FIG. 3, would be
advantageous since Pyrex is highly corrosion resistant, additional
measures must be taken to include the corrosion resistance of the
silicon and other components of the sensing device 12. One method
of improving those structures of the sensing device 12 formed of
silicon, such as the upper and lower cap layers 44 and 46, is to
fabricate those structures of heavily boron-doped silicon. Heavily
boron-doped silicon is believed to be largely corrosion resistant
and/or harmless to tissues in biologic environments.
Another method by which corrosion resistance of the implanted
device 12 may be improved is through coating of the device 12 with
titanium, iridium, Parylene (a biocompatible polymer), or various
other common and/or proprietary thick and thin films. Such a coated
device provides two levels of corrosion resistance: and underlying
stable surface and a separate, stable coating (which may also be
selectively bioactive or bioinert). Provided with these two levels
of corrosion resistance, even if the outer coating contains
pinholes, cracks, or other discontinuities, the device 12 retains a
level of protection.
A number of different, and at times application-specific, schemes
can be envisioned for long-term use of the sensing device 12 of the
present invention. In general, it is necessary to anchor the device
12 so that migration of the device 12 does not occur within the
patient. A dislodged device 12 may migrate away from the
physiologic parameter intended to be sensed, thereby rendering the
device 12 useless for its intended purpose and requiring
implantation of another device 12. A variety of such anchoring
schemes is discussed below.
Referring now to FIG. 14, a screw (or stud) 104 is attached to the
lower cap layer 46 of the sensing device 12. Preferably, the screw
104 is attached to the lower cap layer 46 with biocompatible epoxy
or a similar method. The screw 104 is then embedded into tissue 106
of the patient and the device 12 retained in place. Preferred
materials for the screw 104 include stainless steel and
titanium.
Another scheme for securing the sensing device 12 within a patient
is seen in FIG. 14. As seen therein, the sensing device 12 has
secured to the lower cap layer 46 a sheet of mesh 108. The mesh 108
becomes encapsulated by tissue of the patient over time, thus
anchoring the sensing device 12. Sutures 110 may be used to hold
the sensing device 12 in place until encapsulation occurs.
Preferred materials for the mesh 108 include loosely woven,
biocompatible cloth and the mesh 108 may range in size from 1 to 20
mm.
An endoluminal attachment scheme is illustrated within FIG. 15. In
this application, sensing device 12 is attached to stent-like
spring cage 112. As such, the sensing device 12 may be
non-surgically injected into a blood vessel 114 or other body
cavity containing fluid flow. After ejection from the insertion
apparatus (not shown), the spring cage 112 expands and lodges the
sensing device 12 at the sensing location, while allowing blood (or
other fluid) to continue flowing past the sensing device 12. To
expand outward, the spring cage 12 is formed so that the arms 115
thereof are resiliently biased outward. Preferred materials for the
arms 115 include stainless steel or titanium. The arms 115 may also
be in wire or other forms.
Another endoluminal attachment scheme is shown in FIG. 16. In this
embodiment, the sensing device 12 is anchored in place within
vessel 114 by a set of radially outwardly expandable spring arms
116. The spring arms 116 may be provided with depth-limited
anchoring tips 118 on their ends to further secure the sensing
device 12. The arms 116 may be in wire, ribbon or other form and
are biased outwardly to cause engagement of the anchoring tips 118
with the wall of the vessel 114. Preferred materials for the arms
116 and for the anchoring tips 118 include stainless steel or
titanium.
In FIG. 17, the sensing device 12 is encapsulated in a
biocompatible material such as poly(methyl methacrylate), yielding
a pellet-like profile designated at 120. A recess 122 formed in the
pellet 120 allows access to the movable element 64. In addition to
providing an alternate form factor that may be less mechanically
irritating to tissue 124 both during and after implantation, such
an embodiment may better allow the sensing device 12 to be
incorporated into the body of a medical device, such as an
extrusion, injection-molded part, soft rubber, or other material,
that otherwise would poorly anchor to a rectangular or other
geometrically shaped sensing device 12. Obviously, encapsulation
could be used to give the sensing device other profiles or form
factors as well.
From the above, it can be seen that many applications exist for the
system 10 of the present invention. Some illustrative examples of
such applications are described hereafter.
One application of the described technology, depicted in FIG. 18,
locates the sensing device 12 in an electrode tip 126 of an
implantable stimulation lead 128, such as a stimulation lead used
for cardiac pacing. In such an arrangement, the sensing device 12
could be used with the read-out device 14 for monitoring arterial,
atrial, ventricular, and/or other blood pressures.
In the application seen in FIG. 19, three sensing devices 12 are
being used to calculate a diameter 130 of a flow path 132 defined
by walls 134. In addition to the diameter 130, mass and/or
volumetric blood or other fluid flow rates through the flow path
132 may be calculated. The sensing devices 12 are located in a
variable diameter catheter 136 or similar geometric construction
conductive to taking such measurements. Computational fluid
dynamics (CFD) models and calculations utilizing the distances
between the sensing devices 12 (L.sub.1 and L.sub.2) and pressure
changes .DELTA.P.sub.1 and .DELTA.P.sub.2 therebetween, can be used
to derive the desired parameters from suitably precise pressure
data.
Cardiac monitoring applications can particularly benefit from the
present system 10 in its various embodiments. One possibility is to
locate the sensing devices 12 (either by means of a multiple-sensor
catheter or individually placed sensor devices 12 (or placed as a
tethered pair)) at appropriate locations around a natural or
artificial heart valve or other biologic valve, to monitor the
pressure on either side of, and/or the flow through, the valve. The
same setup may also be used to monitor pressure along a vascular
stent 137, as shown in FIG. 20. Sensing devices 12 may be placed at
one or more locations 138-142 along the length of the stent.
Referring now to FIG. 21, a sensing device 12 is located such that
pressure is measured externally through a vessel wall 144, such as
the wall of a blood vessel. The sensing device 12 is placed in
intimate contact with the wall 144 through use of a variety of
means, including adhesive clips 146 (of a biocompatible material),
tissue growth or other methods. The sensing device 12 is oriented
so that the moveable element 64 is adjacent the vessel wall 144 and
measures pressure transduced through the vessel wall 144. A
calibration factor in active circuitry may be used to adjust the
measured value to an actual value so as to account for the effects
of sensing the pressure through the vessel wall 144.
As an alternative to the foregoing embodiments, the pressure sensor
18 of the sensing device 12 may be augmented and/or replaced with a
structure or sensor 18' that allows a parameter other than pressure
to be sensed. For clarity, in FIG. 22 only the sensor 18' portion
of the sensing device 12 is shown, the nonillustrated elements
being as previously discussed. In the sensor 18', a
chemical-sensitive substance 148 is placed in a confinement cavity
149 and contact with and exterior surface of sensor diaphragm 150.
Osmotic expansion of the substance 148, in response to the
concentration of a target chemical, generates a pressure on the
diaphragm 150 and allowing the concentration of the chemical to be
monitored. For convenience, only the substrate 20 is illustrated,
the fixed electrode and associated structures be omitted. This
sensor 18' may optionally include cap structure 152 to restrict the
expansion of the chemical sensitive substance 148 to the center of
the diaphragm 150 to maximize deflection of the diaphragm 150. A
micromachined mesh, grid, or semipermeable membrane 154, also
optional and either integral to the cap or attached separately
thereto, may be included to prevent the chemical sensitive
substance 148 from escaping (or bulging out of) the confinement
cavity 149, and/or to prevent foreign materials from entering the
cavity 149. The mesh 154 could also exist on the molecular level,
being formed of a material such as a cross-linked polymer.
In another alternative parameter sensing embodiment, a material
with high thermal coefficient of expansion is placed between
moveable and fixed electrodes in a structure otherwise constructed
similar to a capacitive sensor structure, thereby forming a
temperature sensor.
FIG. 23 illustrates an alternative capacitive sensor 156 on the
substrate 20, additional structures are omitted for clarity. In
this sensor 156, the capacitance changes due to a varying
dielectric constant within the capacitive gap defined between
electrodes 158 and 160. The gap is filled with sensing substance
162 chosen such that its dielectric constant changes in response to
the particular physiologic stimulus being evaluated. FIG. 24
depicts an alternate implementation of the above embodiment, with
the electrodes 158' and 160' and the sensing substance 162 being
stacked vertically on the substrate 20, as opposed to the lateral
orientation in FIG. 23.
The pressure, temperature or other data sensing technology, in its
various forms, may be incorporated into an open or closed-loop
therapeutic system for the treatment of medical conditions which
require or benefit from regular, subcutaneous monitoring of
pressures or other parameters. The system may be used, for example,
to control the administration of drugs. One particular application
of this would be to control hyper- or hypotension. In the preferred
embodiment, pressure data from the sensor, alone or in conjunction
with other real-time or preexisting data, is used to adjust drug or
other therapy for hypo- or hypertensive patient. Therapy is
provided by means of a control module worn by, or implanted within,
the patient (similar to e.g., an insulin pump for diabetics). The
module may alert the user to take action, directly administer a
drug intravenously, and/or initiate other invasive or non-invasive
responses. Furthermore, relevant information (including, but not
limited to, measure physiologic parameters, treatment regimens,
data histories, drug reservoir levels) can further be transmitted
from the control module to other locations via cellular phone,
wireless infrared communication protocols or other communication
methods and mechanisms.
Other applications of the implantable wireless sensing device of
this invention include, without limitation, the following: 1)
Monitoring congestive heart failure patients such as left ventricle
pressure monitoring, left atrium pressure monitoring and pulmonary
artery pressure monitoring; 2) other hemodynamics parameters
including blood pressure, blood flow velocity, blood flow volume
and blood temperature; 3) diabetic applications including glucose
level monitoring; 4) urinary applications such as bladder pressure
and urinary tract pressure measuring; and 5) other blood parameters
including O.sub.2 saturation, pH, CO.sub.2 saturation, temperature,
bicarbonate, glucose, creatine, hematocirt, potassium, sodium,
chloride; and 6) cardiac parameters including (previously
discussed) valve pressure gradients and stent pressure
gradients.
In addition to single sensor, an array of different sensors may be
fabricated or assembled on one sensing device to enhance artifact
removal and/or selectivity/differentiation between signals. A
discussion of such a construction best details this construction.
Local pressure or pH variations can add spurious signals to a
pressure- or pH-based glucose sensor. To compensate for these
spurious signals, adjacent pH or pressure reference sensors may be
implemented to measure these environmental parameters. External
sensors may also be used to compensate for factors such as
atmospheric pressure. A combination of sensor arrays, fuzzy logic,
look-up tables, and/or other signal-processing technologies could
all be used to effect such compensation.
The foregoing disclosure is the best mode devised by the inventor
for practicing the invention. It is apparent, however, that several
variations in accordance with the present invention may be
conceivable to one of ordinary skill in the relevant art. Inasmuch
as the foregoing disclosure is intended to enable such person to
practice the instant invention, it should not be construed to be
limited thereby, but should be construed to include such
aforementioned variations, and should be limited only by the spirit
and scope of the following claims.
* * * * *